As shown in FIG. 1A, a conventional biosensor system 1000 typically includes an array of biosensors 1010 for interaction with a buffer solution 1001 that contains the biomaterial under test. The system further includes signal preamplifiers and driving circuits 1020, which are disposed outside the solution 1001 and are connected to the respective biosensors 1010. The signal preamplifiers and driving circuits 1020 are connected to a data analysis and storage unit, such as a computer 1030.
A typical array of biosensors 1010 is shown in cross-section in FIG. 1B, in which a plurality of gold electrodes 1011 are formed on a glass substrate 1012. Typically, these electrodes are connected to the signal preamplifiers and driving circuits 1020 by using suitable connectors or via a set of probing needles. To prevent short circuit, the glass substrate 1012 and gold electrodes 1011 are passivated using a suitable passivation material 1016, with openings over portions of the gold electrodes 1011. Different types of bioreceptive layer 1015 (commonly known as a probe) are formed covering each of the gold electrodes where exposed by the openings 1011 to interact with the biomaterial under test (commonly known as a target). Specifically, the bioreceptive layer may be a phage or enzyme that binds with a predetermined DNA or RNA strand, a peptide or another biological molecule, thereby changing the resistance and/or the capacitance of the circuit in which the electrode is connected. These changes can be detected in turn to establish whether the predetermined biological molecule is present, and even its concentration. The reliability of detection can be enhanced by increasing the number of different bioreceptive layers, which allows cross-referencing and cross-elimination in the data analysis process.
The use of gold for the electrodes 1011 enhances the adhesion of the materials used to form the bioreceptive layer 1015, and the electrodes 1011 are spaced at a sufficient separation to reduce the effects of cross-talk.
On the left side of FIG. 1A, the array of biosensors is disposed in a flow cell 1040, in which the solution or sample 1001 to be analysed is passed over the array, in a direction indicated by the arrow. On the right side of FIG. 1A, the array of biosensors is submerged in a container, such as a test tube 1041, in which the solution or sample 1001 containing biomaterial under test is held.
Although generally found to be effective, there are commonly problems in such arrangements in that the comparatively long signal path due to the wiring between the electrodes and the signal pre-amplification and driving circuits 1020 picks up noise, thereby reducing the signal-to-noise ratio and greatly limiting the system sensitivity.
This problem has been mitigated at least to an extent by integrating the pre-amplification and driving circuitry with the sensors on the chip. A schematic representation of this is shown in cross-section in FIG. 1C. This figure is the same as FIG. 1B in most respects and like reference numerals indicate like parts. However, FIG. 1C further shows a Thin Film Transistor (TFT) chip 1050 provided on the glass substrate 1012, the illustrated TFT comprising a gate 1051, source 1052, drain 1053 and channel 1054. Of course, a plurality of TFTs can be provided on the sensor chip and interconnected to provide the desired pre-amplification and driving operations. In addition, other forms of transistor technology, such as CMOS, are possible.
The integration of the pre-amplification and driving circuitry on the chip has the significant advantage of an increase in signal-to-noise ratio. Therefore, for a fixed concentration of sample solution, this integration will require smaller probe areas, which is beneficial in quality control applications. The sensitivity of the probe is also improved, which is beneficial for medical applications when the volume of sample is limited in supply. However, as shown in FIG. 1C, this implementation still requires external electronics 1060 for the detection and relay of signals from the sensor chip to the computer 1030. This is a significant obstacle in terms of size and set up cost.
In the implementations shown in both FIGS. 1B and 1C, the pre-amplification and driving circuits to which the electrode is connected is typically a potentiostat. A physical representation of a prior art potentiostat circuit is shown in FIG. 2A. and an electrical representation is shown in FIG. 2B. As shown in FIG. 2A, the potentiostat circuit consists of three electrodes: a counter electrode CE, a reference electrode RE and a working electrode WE. The electrodes are connected to an electrochemical cell, which is realised by a buffer solution with the biomaterial 1001 between counter electrode CE and reference electrode RE, and a bioreceptive layer 1015 between reference electrode RE and working electrode WE. An equivalent circuit of the three terminal electrochemical cell is given in FIG. 2B. In essence, the potentiostat circuit measures the current at working electrode WE due to a redox reaction at the bioreceptive layer while RE is kept at a known voltage. Counter electrode CE is provided to supply the necessary current to maintain reference electrode RE at a known voltage. Many versions of potentiostat are known and a multi-channel potentiostat is known for the measurement of a matrix of samples. The applied input voltage changes the redox (reduction-oxidation) reaction at the bioreceptive layer, which in turn changes the effective values of the capacitance and resistance of the equivalent circuit, allowing the detection of the presence of a target biological molecule and its concentration.
Cyclic voltammetry is a technique commonly employed in a potentiostatic measurement. Here, the voltage at RE, shown as VSCAN in FIG. 2B, is a triangular wave. During the first half cycle, a reduction (or oxidation) reaction is dominant, diffusion occurs due to different concentration of the reduced and oxidised species, and the net current flow is reflected as a peak at the output. During the second half cycle, if the redox reaction is reversible, an oxidation (or reduction) reaction occurs, and a peak current of the same magnitude but in the opposite direction can be seen at the output. A typical output for a reversible redox reaction is given in FIG. 2C. It is of importance to note the height and locations of any peaks in such an output.
Some redox reactions can be particularly slow and therefore time-consuming to complete a measurement. The measurement of a large number of redox reactions for different target biological molecules under different conditions can therefore be highly problematic.
However, the present invention is not limited to potentiostat circuits. Other means of measuring or sensing properties of a sample can also be used. One such means is a pH sensor, which typically includes an ion-sensitive field effect transistor (ISFET) with an exposed gate in contact with the sample solution, the drain-source current of which is found to be related to the concentration of H+ ions in the solution. Since the concentration of H+ ions is indicative of pH, the pH of the sample solution can be measured.
The use of RFID (radio frequency identification) circuits in biosensor chips is also known. This removes the requirement for a physical interconnection between the sensor chips 1010 and the data analysis and storage computer 1030. However, problems still remain with the speed and throughput with which samples of solutions can be measured, which leads to increased costs and backlogs in hospitals and other laboratories, together with the problems concomitant with such backlogs. This is a particular problem with measurements of redox reactions of bioreagents, which tend to be slow. It is therefore very time-consuming to complete all measurements in an array of electrodes.
RFID devices are either active, wherein the device includes a power supply, or passive, wherein the device extracts power from received radio signals and uses this to power itself.